The present invention relates to the art of electromagnetic shielding. It finds particular application in conjunction with gradient coils for magnetic resonance imaging apparatus and will be described with particular reference thereto. However, it is to be appreciated that the present invention also finds application in conjunction with systems which employ gradient magnetic fields and other applications in which electromagnetic shielding is desired.
In a magnetic resonance imaging system, gradient coil assemblies are commonly pulsed with electrical current pulses to produce magnetic gradients across the main magnetic field in the vicinity of an imaging region. As an unwanted side effect, magnetic field gradients are produced which interact with external metallic structures such as the magnet cold shields, the magnet dewar, and the like. The interaction generates eddy currents in the effected structures. These eddy currents, in turn, generate eddy magnetic fields which have a deleterious effect on the temporal and spatial quality of the magnetic field in the vicinity of the imaging region and, hence, in the resultant image quality.
The eddy current problem is often addressed by placing an active shielding coil between the primary gradient coil and the effected structure. The shielding coils are designed to substantially zero or cancel the magnetic field external to itself thereby preventing the formation of eddy currents in potentially vulnerable structures.
Previously methods for production for magnetic gradients in magnetic resonance imaging systems consisted of winding discrete coils in a bunched or distributed fashion on an electrically insulating hollow cylindrical former and driving the coils with a current source of limited voltage. Conventional bunched coil designs include the Maxwell and the Modified Maxwell Pair for z-gradient production, and the Golay or Modified Golay (multi-arc) Saddle Coils for x and/or y-gradient production. Typically, these methods consisted of iteratively placing coil loops or arcs on the cylindrical former until the desired gradient strength, gradient uniformity, and inductance (related to stored energy) were achieved. These previous designs were generally developed in a "forward approach" whereby a set of initial coil positions were defined (i.e., the initial coil distribution), the fields and the inductance/energy calculated, and if not within particular design parameters, the coil positions would be shifted (statistically or otherwise) and results re-evaluated. The iterative procedure continued until a suitable design was obtained.
More recent methods of generating magnetic fields in magnetic resonance imaging systems utilize an "inverse approach." In the "inverse approach" method, the gradient magnetic field is forced to match predetermined values at specified spatial locations inside the imaging volume and a continuous current density is calculated which is capable of producing such a field. The "inverse approach" method assumes that the primary gradient coil has finite dimensions while those of the secondary or shield coil are left unrestricted (infinite). After the generation of continuous current distributions for both the primary and the shield coils, an apodization algorithm is performed on the continuous current density of the shield coil in order to restrain it to desirable dimensions. Following the modification of the shielding coil's continuous current, the Stream Function technique is employed in order to obtain discrete current patterns for both coils. Application of the Biot-Savart law to the discrete current pattern ensures that the discretization procedure was proper. This approach created generally more energy efficient gradient coil assemblies with higher gradient strengths and faster slew rates as compared to the "forward approach" method.
In certain applications, for example imaging of the head, it is desirable to maintain the sweet spot of the gradient coil assembly at a predetermined location corresponding to a particular anatomy of interest. However, due to a subject's surrounding anatomy (the subject's shoulders in this case) the gradient coil assembly can not be appropriately centered. That is to say, when the gradient coil assembly's diameter is large enough to encompass a subject's shoulders it is too far radially from the subject's head, and conversely, when the diameter is sufficiently close to the head, the shoulders interfere with an end of the coil such that the isocenter of the gradient coil is not properly aligned with the area of interest in the head. Shortening the length of the gradient coil so that the shoulders of a subject do not come into play reduces the usable imaging volume and deteriorates linearity. Attempts to adjust the sweet spot of the coil by means of a non-symmetric gradient coil assembly runs afoul of high torques generated by the interaction of the main magnetic field with the currents of the gradient coil assembly and less desirable gradient fields. One approach used to address the issue, was to employ a symmetrical cylindrical gradient coil assembly with flared ends that could accommodate a subject's shoulders. Added benefits of the flared design are increased patient access and reduced claustrophobia. However, heretofore shielding for this flared design has been lacking.
In one particular approach discussed in U.S. Pat. No. 5,497,089 to Lampman et al., an insertable cylindrically shaped gradient coil assembly is presented with flared ends. However, the gradient coil assembly is not shielded. In this case, eddy current effects not suppressed by shielding prohibit the use of fast imaging sequences that are particularly sensitive thereto.
Another particular gradient coil assembly is presented in a paper by Schenck et al., "Design Criteria for a Folded Gradient Coil", 5th ISMRM (Vancouver, Canada), April 1997, pg. 1468. The paper presents a design methodology for a flared shielded gradient coil using the "forward approach". The design model presented assumed a series connection between the primary and shield coils such that they share the same current. This design resulted in inadequate shielding as compared to a traditional shielded design.
U.S. Pat. No. 5,378,989 to Barber et al. presents yet another flared gradient coil assembly for use with open magnet systems. However, the flared portions are not at the ends of a cylindrical coil. Rather, the flared portions are near the isocenter. This prohibits the assembly from being used with a closed bore magnet. Furthermore, the design restricts the flare to a 90.degree. angle and is based on a "forward approach" which results in decreased performance.
For interventional procedures and like applications where patient access is desirable, it is advantageous to design the gradient shielding coil such that its dimensions do not exceed those of the flared primary gradient coil. In this manner, patient access can be maximized and the feeling of openness can reduce patient claustrophobia. However, in general, previous methods and prior art suffer the drawback that as the shielding coil length approaches that of the primary coil, increased levels of eddy current effects within the imaging region deteriorate image quality. Conversely, when sufficient shielding is achieved, the dimensions of the shielding coil are substantially larger than those of the primary coil such that the level of patient access is encumbered and an increased level of patient claustrophobia is experienced.
The present invention contemplates a new and improved shielded gradient coil assembly and method for designing such which overcomes the above-referenced problems and others.